1. Field of the Invention
This invention relates generally to the diagnostic testing of a body fluid, especially blood, and more particularly, to disposable sensor assemblies for use in automated bedside monitors to measure the glucose content of body fluid.
2. Description of the Related Art
Presently employed electrochemical glucose sensors are essentially modified polarographic oxygen electrode assemblies. They are two or three-electrode electrochemical cells covered by surface layers of an immobilized enzyme, e.g. glucose oxidase. The glucose oxidase layer, for example, not only serves as an electrolyte but also catalyzes the reaction of glucose with oxygen to form gluconic acid and hydrogen peroxide, H.sub.2 O.sub.2. EQU glucose+O.sub.2 .fwdarw.gluconic acid+H.sub.2 O.sub.2
Thus, the glucose level can be quantitatively determined by either measuring the reduction in the oxygen partial pressure, pO.sub.2, or increase in H.sub.2 O.sub.2 by means of the underlying polarographic oxygen electrode assembly. When the glucose sensor operates in the oxygen mode, the working electrode is negatively polarized against a counter electrode to measure the limiting oxygen reduction current, which is proportional to the pO.sub.2. Without the enzyme coating, the sensor is a polarographic oxygen electrode. Conversely, the polarity of the applied bias may be reversed and set at another level to measure the anodic current produced by the oxidation of hydrogen peroxide. There are advantages and disadvantages in either mode. If one chooses to use pO.sub.2 as an indirect measurement of glucose, then the oxygen partial pressure of the sample must be determined. If hydrogen peroxide is measured, it is no longer necessary to know the pO.sub.2 of the sample. However, other compounds in the sample, such as ascorbic acid, may also be oxidized at the same potential as the hydrogen peroxide, H.sub.2 O.sub.2. These other components can thus produce an interference effect, and render the determination of glucose inaccurate.
In either mode of operation, the glucose sensor must deal with a common problem; that is, in undiluted whole blood, there is a very large excess of glucose relative to oxygen and the reaction of glucose with oxygen becomes oxygen limited. In this case, the sensor is saturated and sensor saturation thus means that the glucose sensor output approaches zero because there is no more oxygen available on the sensor surface, and that the sensor is therefore glucose saturated. In other words, the range of glucose concentrations detectable by such a sensor is very narrow, typically only up to 40 milligrams per deciliter (mg/dl) of blood before reaching saturation. Clinically, a linear response up to 500 mg/dl is required.
Several approaches have been used to expand the useful range of such glucose sensors. For example, in a commercial glucose analyzer manufactured by Yellow Spring Instrument Inc., as disclosed in U.S. Pat. No. 3,539,455 issued to Leland Clark, whole blood samples are diluted prior to each measurement. Yet another approach is to design the sensor in such a way that oxygen is readily available to the enzyme layer across a hydrophobic structure while the access of glucose to the enzyme layer is restricted through a long and narrow hydrophillic path such as those schemes described in U.S. Pat. Nos. 4,650,547 and 4,890,620. All these approaches require the reduction of the molar glucose to oxygen ratio to extend the sensor range by relatively sophisticated means. Another approach is to put a semipermeable membrane between the enzyme layer and the whole blood sample such as those described in U.S. Pat. Nos. 4,757,022 and 4,759,828. Oxygen can diffuse across this semipermeable membrane much more easily than can glucose, typically at a ratio of more than 100 to 1. This membrane drastically slows down the diffusion rate of glucose and thus creates a more favorable glucose to oxygen molar ratio to prevent premature saturation of the sensor. The present invention eliminates the need for such a highly selective semipermeable membrane and thus allows the direct use of a fairly simple enzymatic glucose sensor in a dynamic mode with or without a semipermeable membrane. 3. Summary of the Invention
The present invention provides a method of using a simple glucose oxidase-based electrochemical sensor to measure the glucose concentration in biological fluids, especially body fluid, particularly human blood. In this present invention, the sensor is initially exposed to a glucose-free or low-glucose, oxygen-containing solution to register its oxygen reduction limiting current, also known as the zero-glucose or low-glucose baseline. Then, a sample is allowed to come in contact with the sensor for a certain amount of time, such as 30 seconds, to allow glucose to diffuse into the enzyme layer until saturation occurs. Then, the sensor is removed from the sample and again placed in the glucose-free, or low-glucose, oxygen-containing solution to allow glucose trapped in the enzyme layer to diffuse away. Once the diffusion process is completed, the sensor output will return to its original zero-glucose, or low-glucose, baseline. Since the amount of glucose trapped in the enzyme layer is a function of the glucose concentration in the sample, the time it takes to return to the full zero-glucose, or low-glucose, baseline output, or a fraction of it, from the time the sensor is returned to the glucose-free, or low-glucose, solution (defined as time-to-recover) has now been discovered to be a function of the original glucose concentration in the sample alone if other conditions, especially sample exposure time and temperature, are held constant. By choosing an enzyme layer of a suitable thickness, sufficient resolution in time-to-recover enables us to accurately measure samples ranging in glucose concentration from approximately 50 to 600 milligrams/deciliter with a single sensor.
In an attempt to further improve the resolution of the glucose sensor, a very thin membrane layer of a barrier material can be applied over the enzyme layer primarily to slow down the diffusion of glucose into and out of the sensor and in particular to slow down the diffusion of glucose to the immobilized enzyme layer of the sensor. This barrier membrane layer is typically less than 0.001 inches thick but may also be less than 0.010 inches thick and may be a polycarbonate film or a perfluorinated ionomer membrane such as that ionomer membrane sold under the trade name Nafion. This barrier membrane layer typically has a low diffusion coefficient for glucose and high diffusion coefficient for oxygen and must in general be more than 0.00001 inches thick. It not only slows down glucose diffusion out of the enzyme layer for better resolution but also permits oxygen in the baseline solution to reach the sensing electrode in time for the time-to-recover measurement. However, in its operation here it acts differently from the conventional glucose sensor approach described in 4,650,574 which requires the barrier to maintain a glucose to oxygen ratio of at least 1 to 100 to avoid sensor signal saturation over the range of clinical interest, 50-600 mg/dl. This present invention works even at a ratio 1 to 10 or below because it does not quantitatively determine glucose concentration based on the measurement oxygen current directly. Therefore, it is no longer critical to avoid sensor saturation due to lack of oxygen in the enzyme layer. In addition, unlike conventional enzymatic glucose sensors operating in the oxygen reduction mode, this present invention does not require a knowledge of pO.sub.2 in the sample to compensate for an initial offset in the event that pO.sub.2 values may be expected to vary over a wide range in the case of critically ill patients.
The specific form of the time-to-recover versus sample glucose concentration function depends on the structure of the sensor. In practice, an empirical calibration curve can be obtained to translate recovery time into glucose concentration.